Sensing platform for quantum transduction of chemical information

ABSTRACT

A system for determining chemistry of a molecule in a high background interfering liquid environment by application of an electronic signal at a biased metal-electrolyte interface is disclosed. One or more of a resonant exchange of energy between one or more electrons exchanged by the metal and the electrolyte and vibrating bonds of a molecular analyte, for example, may be sensed by measuring small signal conductivity of an electrochemical interface.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims the benefit of U.S. Provisional Application No.61/681,380, filed Aug. 9, 2012, and is a continuation of U.S.Non-Provisional application Ser. No. 13/963,272, filed Aug. 9, 2013, theentire contents of both of which are hereby incorporated herein byreference.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

This invention was made with government support under contractN66001-11-1-4111 awarded by the Defense Advanced Research ProjectsAgency. The government has certain rights in this invention.

BACKGROUND

Identification and analysis of chemical and molecular species within anenvironment is well established. Typically, electronic systems rely uponalterations in current, voltage, or charge to indirectly qualify andquantify chemical analytes. Bioassays detect analytes indirectly bymeasuring various molecular interactions. Some bioassays measureanalytes by activating a label that is covalently attached to a bindingpartner, upon analyte binding to a bait molecule. Other bioassaysmeasure analyte binding to an immobilized bait molecule to a solidsubstrate and measuring changes in charge, refractive index, or masschange at an interface between the solid substrate and liquid sample.Demand for a low-cost and field-use friendly method of low concentrationanalytes has resulted in ongoing efforts to improve the functionalityand practicality of chemical and molecular detecting devices.

BRIEF DESCRIPTION OF THE DRAWINGS

Many aspects of the disclosure can be better understood with referenceto the following drawings. The components in the drawings are notnecessarily to scale, emphasis instead being placed upon clearlyillustrating the principles of the present disclosure. Moreover, in thedrawings, like reference numerals designate corresponding partsthroughout the several views.

FIG. 1A shows transition states of a nonadiabatic reaction.

FIG. 1B shows transitions states of an adiabatic reaction.

FIG. 2A shows weak coupling between initial and final electronic energystates in the nonadiabatic reaction of FIG. 1A.

FIG. 2B shows strong coupling between initial and final electronicenergy states in the adiabatic reaction of FIG. 1B.

FIG. 3 is a schematic depicting a measurement of a flux of electronscrossing an electrified dielectric monolayer modified electrochemicalinterface of a biosensor.

FIG. 4 shows factors affecting energy state transition rate in a weaklycoupled, non-adiabatic reaction.

FIG. 5 shows data suggesting sensitivity of the biosensor to a singleatom isotope substitution.

FIG. 6 shows an exemplary embodiment of the biosensor in which a sensordie with a patterned electrochemical interface and a CMOS interface chipare integrated into a low-cost, disposable, lateral flow-basedmicrofludic architecture.

FIG. 7 shows disposable modules that make up a fluidic system for oneembodiment of the biosensor.

FIG. 8 shows an enlarged view of a sensor package with electrode sensorsarranged in an array.

FIG. 9 shows steps in generating a reference database.

FIG. 10 shows tunneling barriers at metal-dielectric anddielectric-electrolyte interfaces.

FIG. 11 shows sequential layering of high and low k-dielectric materialsfor a high-k dielectric insulator.

FIG. 12 shows a magnetic tunneling film architecture that usesdifferentially oriented film magnetic moments to further restrictelectronic transition.

FIG. 13 shows an example of three-electrode feedback suppression ofthermal noise for electronic transition measurements.

FIG. 14 shows a table containing program metrics.

DETAILED DESCRIPTION OF THE EMBODIMENTS

Reference is now made in detail to the description of the embodiments asillustrated in the drawings. While several embodiments are described inconnections with these drawings, the disclosure is not intended to belimited to the embodiment or embodiments disclosed herein. On thecontrary, the intent is to cover all alternatives, modifications, andequivalents.

For a variety of applications, detection and identification of smallamounts of various molecules is desired. Typical molecular species whosedetection is desired include, but is not limited to, small moleculeanalytes like amino acids and metallic ions to large biological likeproteins, DNA, and RNA. In particular, detection of biomarkers inbiological samples is important for disease detection, disease analysis,and disease pathway investigation. Further, detection of contaminants inenvironmental samples, such as water, is important for homelandsecurity, public safety, and environmental welfare.

For example, an ideal platform for use in detecting biological threatsshould be well suited to identifying a large range of harmful agents andtoxins. As many of these agents and toxins are highly infective, theplatform must demonstrate great sensitivity and specificity to allowearly exposure detection, reduce false positives, and enable targetedcountermeasures and minimize spread of infection. The platform must alsoallow for rapid detection to enable timely intervention. The challengeof developing a sensitive, yet specific, high throughput detector with alarge working range may be appreciated. The challenges are complicatedby a need for the detector to be portable and have minimal operationalcomplexity, low power consumption, low manufacturing cost, andoperability in harsh environments.

Platforms for detecting molecules in samples have evolved over theyears, from purely laboratory based and impractical for point-of-usedetection, to portable miniaturized “Lab-on-a-Chip” platforms. Forexample, detection of biological threats has evolved from conductingthreat detection and diagnosis though Laboratory Response Network, to amobile lab based system, such as Biological Integrated Detection System(BIDS), to mesoscale peptide bioassays. This evolution is a directreflection of the need for small molecule detectors that are capable ofrapid and point-of-use detection.

Traditional bioassays fall into two categories: label-based orlabel-free. In label-based bioassays, the target molecule, such as atoxin, binds with a bait molecule, often a complementary peptide, DNA,or RNA molecule, which has a covalently attached label. Fluorescent dyesand radioactive isotopes are commonly used labels where binding of thetarget molecule to the bait molecule causes the release of fluorescenceor radiation. Measurement of fluorescence or radioactivity provides anindirect detection and quantification of the target molecule.

These array label-based assays suffer from significant limitationsdespite improved sensitivity and specificity. First, these arraylabel-based systems require identification, design, synthesis andimmobilization of the bait molecules, which are significantlyrate-limiting in the assay manufacturing process. Second, immobilizationof a bait molecule with a complex three-dimensional structure results ina loss of activity of the bait molecule, which generates false negativeoutcomes. Third, the addition of a covalently bound fluorophore orradioactive tag significantly modifies an interaction between the targetmolecule and the bait molecule, resulting in false positives andnegatives. Fourth, tagging a bait molecule with a fluorescing orradioactive label adds a layer of complexity to the manufacturingprocess. Fifth, the assay requires that readers to detect theoptical/radiation signal from the tags be incorporated with theplatform, thus dramatically increasing platform cost and reducingplatform portability. Finally, extinction of a signal generated from abinding event due to scattering from the background matrix is apersistent problem resulting in low signal-to-noise.

The limitations imposed by traditional labeled bioassays prompted thedevelopment of the label free methods. Like the label-based assay, thelabel free assay is array based with bait molecules immobilized on asolid substrate. The detection of the binding between the targetmolecule and bait molecule is based on (a) the change in charge at thesolid-liquid interface that results from the binding event, (b)evanescent waves attenuation due to a change in refractive index at thesolid-liquid interface, or (c) mass change at the solid-liquidinterface. Charge based detection methods eliminate the need forexpensive signal readers, thereby reducing the cost of detection,enhancing system portability, reducing overall power consumption, andincreasing ease of operation. The charge based method is also scalable,which is an essential strategy in developing a high throughput detectionplatform. Though the label-free platforms do not suffer from problemslike tag-altered target molecule binding and reduced signal yield, theyare still afflicted by the issue of bait molecule misfolding onimmobilization to a solid surface.

The bait molecule is utilized to infer whether the target molecule ispresent or absent in both label-based and label-free platforms. Theactual identity of the target molecule is inferred from the nature ofthe bait molecule with which binding occurs. Mass spectrometry, on theother hand, is a time-critical, broadband analysis technique thatdirectly measures molecular composition from estimates of charge-to-massratios of vaporized fragments of the analyte. Commercial massspectrometers are reportedly capable of detection in the nanomolarconcentration range. Arrayed, multi-channel, modular architectures fortime-of-flight (TOF) mass spectrometers have been detailed for rapid,in-parallel acquisition of information.

However, mass spectrometry analysis is better suited to larger molecularweight target molecules that can be fragmented into several constituentmoieties for analysis. Small molecular weight (<5 kDa) target moleculesare not easily identified by this technique. The mass spectrometer andassociated ancillary equipment (like vacuum pumps) are energy intensivein operation and are not easily miniaturized, thus making portability anissue. Additionally, mass spectrometer operation and data analysisrequire intervention of skilled technicians, making the detectionplatform ill-suited for point-of-use applications. Thus, in view oftraditional detection systems, one of ordinary skill in the art willappreciate the need for a robust, rapid, low-cost, point-of-usedetection platform for small amounts of molecules in fluid samples.

Molecular vibration-assisted-charge transfer between electron source inresponse to a molecule has been documented in nature. Fruit flies detectodorants by transferring an electron from an intracellular electronsource upon entrance of an odorant into a transmembrane pocket. Theelectron charge transfer stimulates G-protein mediated signaltransduction pathways and thus allows the fruit fly to identify anodorant. Similarly, according to the embodiments described herein,detection of molecular analytes by detection of an electron transfer isachieved. In the biosensor according to the embodiments describedherein, current measured due to electron transfer that containsinformation about vibrational frequencies of molecular bond vibrationswithin a molecular analyte is acquired directly from a physicaltransducing interface and analyzed thereafter.

Generally, the biosensor system according to the embodiments describedherein comprises an electrochemical charge transfer platform configuredto be “slow”, relative to a speed of molecular vibrations of moleculesin a liquid sample. This electrochemical charge transfer platformcomprises a metal layer and an electrochemical interface separated by adielectric layer. The electrochemical interface interacts with anelectrolyte, or diffuse, layer, which comprises the liquid sample. Thedielectric layer acts as a molecular insulator to slow down the rate ofelectron transfer. The biosensor system further comprises a high gainnoise suppression feedback loop to electronically “cool” the system andminimize thermal noise that may otherwise destroy a signal of interest.At low electronic temperatures, transfer of electronic charge occurs ina resonant manner by quantized vibrations of a target analyte. Thesystem measures resonant interactions by measuring small signalconductance across the electrochemical interface. Each resonance,detected on a conductance profile, is correlated to a vibrationalfrequency of a molecular bond in the analyte. As vibrational frequenciesare characteristic signatures of molecular bonds, akin to humanfingerprints, the number and types of bonds in the analyte can bedetermined from a conductance profile. Each analyte possesses a uniquemolecular bond signature, thus allowing direct, highly specific analytedetection.

In one embodiment, a system for sensing chemical information isdescribed. The system includes a sample acquisition zone, a filtrationmodule operatively coupled (e.g., via wicking, etc.) to the sampleacquisition zone, an immunoseparation module operatively coupled to thefiltration module, a tapered micro-chromatogram operatively coupled tothe immunoseparation module, and an adsorption pad operatively coupledto the tapered micro-chromatogram. In one embodiment, the system furtherincludes a quantum tunneling biosensor interface mounted on a shieldedprinted circuit board, the quantum tunneling biosensor interface beingoperatively coupled to the adsorption pad. The quantum tunnelingbiosensor interface may include a transducing electrode array includingdielectric thin films deposited on a metal electrode array. The metalelectrode array may be layered on a silicon die, and the silicon die mayinclude through-silicon vias. The biosensor interface may furtherinclude processing logic operatively coupled to the through-siliconvias.

In another embodiment, a system including a quantum tunneling biosensorinterface, a transducing electrode array, and processing logic isdescribed. According to certain aspects, the transducing electrode arraymay be located on or adjacent to the quantum tunneling biosensorinterface. The transducing electrode array includes dielectric thinfilms layered on a metal electrode array. The metal electrode array maybe mounted on a silicon die. The processing logic may be operativelycoupled to the transducing electrode array by through-silicon vias inthe silicon die.

The system may further include a modular fluidic system. The modularfluidic system may include a sample acquisition zone, a coarsefiltration module operatively coupled (e.g., via wicking, etc.) to thesample acquisition zone, an immunoseparation module operatively coupledto the coarse filtration module, a tapered micro-chromatographoperatively coupled to the immunoseparation module, and an adsorptionpad operatively coupled to the quantum tunneling biosensor interface.

According to another embodiment, a method of identifying a targetanalyte in a biological fluid is described. According to variousaspects, the method includes applying biological fluid to a dielectricmonolayer modified electrochemical interface, applying a voltage biasacross the electrochemical interface, and measuring a flux of electroniccharge across the electrochemical interface. In certain embodiments,applying a voltage bias across the electrochemical interface may includetuning the voltage bias to achieve a weakly coupled, non-adiabaticelectronic transfer across the electrochemical interface. The method mayfurther include filtering the biological fluid prior to applying thebiological fluid to the dielectric monolayer modified electrochemicalinterface, wherein filtering the biological fluid includes passing thebiological fluid through a porous membrane resulting in sizefractionated fluid.

According to additional aspects, the method may further include passingthe size fractionated fluid through a membrane coated with a proteinspecific antibody resulting in size fractionated and immunoseparatedfluid. The method may further include passing the size fractionated andimmunoseparated fluid through a micro-chromatograph.

Resonant Electron Transfer at an Electrochemical Interface

The biosensor system according to the embodiments described hereinrelies upon measuring electron flux or leakage currents produced innon-adiabatic, charge-transfer-related transitions at an electrochemicalinterface. The leakage currents provide or are representative ofhigh-resolution molecular structural information. The molecularstructural information, one determined, is unique to each analyte, thusallowing for highly specific molecular species determination.

FIG. 1A shows transition states of a nonadiabatic reaction, and FIG. 1Bshows transitions states of an adiabatic reaction. Charge transferacross an electrified electrode-insulator-electrolyte interface islimited by a quantum-mechanical, electron transition process. The natureof the electron transition and the magnitude of the transition chargeflux depends on the extent of coupling between the initial and finalstates of the transferring electron, which can be non-adiabatic, asillustrated in FIG. 1A, or adiabatic, as illustrated in FIG. 1B.Further, a coupling potential energy is a function of a voltage biasapplied across the electrochemical interface. Coupling strength can betuned by changing an applied voltage bias, for example, by tuning alocal interface chemistry, by conditioning the system to reduceintrinsic noise, or by scaling down a physical sensor of the system.Coupling potential is determined from electrostatic interactions betweentransitioning electrons and ionic charges in a double layer at theelectrochemical interface, as well as on the electrostatic interactionsthat occur with a liquid sample bath which cause thermal dephasing ofresonance phenomena in the charge transfer process.

FIG. 2A shows weak coupling between initial and final electronic energystates in the non-adiabatic reaction of FIG. 1A, and FIG. 2B showsstrong coupling between initial and final electronic energy states inthe adiabatic reaction of FIG. 1B. The coupling between initial andfinal states can be weak, as illustrated in FIG. 2A, or strong, asillustrated in FIG. 2B, and the coupling strength may be tuned, forexample, by applied bias, interface chemistry, interface size andintrinsic interface noise. Where an applied voltage bias allows for anelectron transition and initial and final electron energy states aresignificantly coupled to one another, de-phasing is strong (See FIG.2B). In a strongly coupled electron transfer, the electron wavefunctionis localized to initial and final energy states before and after thecharge transition, which results in particle-like behavior and anadiabatic charge transfer. Thus, the charge transfer is “fast” andlimited only by the rate of dielectric polarization around a reactantand product species. The transitioning electron is always in a groundstate resulting in no possibility for resonant electron transfer tooccur.

In contrast, where the applied voltage bias allows for the initial andfinal energy state of the transferring electron to be weakly coupled toone another in a non-adiabatic reaction, then electron transfer viaelectron tunneling occurs (See FIG. 2A). In this transition, an electronmay be spread over the electrochemical interface as a delocalized waveboth before and after energy state transition. Thus, electron transferis limited by a rate of electron transition from reactant to productstate, which is a function of the composition of the interveningdielectric layer. The transition event in this case is tunnelinglimited. This weakly coupled transition allows a transferring electronto be excited to a higher energy level, unlike the case whenelectrostatic coupling is strong. This allows for the possibility of anelectron transfer that can be resonant (in energy) with molecularvibrational modes between the electrolyte (and analyte contained within)and the metal electrode.

Vibrational Mode Information Transduction Interface

The measurement of the flux of electrons (“leakage” currents) crossingan electrified dielectric monolayer modified electrochemical interfaceallows for analyte detection in the embodiments of the biosensordescribed herein. In that context, FIG. 3 is a schematic depicting ameasurement of a flux of electrons crossing an electrified dielectricmonolayer modified electrochemical interface of a biosensor. Leakagecurrents at electrochemical interfaces have been long ignored throughdecades of focus on traditional charge-based biosensors. A voltage biaswithin the weakly coupled bias window is applied to the dielectricmonolayer modified electrochemical interface. This results in electronictransfer with a diffuse electron spread across the dielectric monolayerand a resonant energy exchange between quantized energy of theelectron-wave and the molecular vibrational modes of analytes in theelectrolyte (diffuse) layer. The electron flux or leakage current ismeasured as an impedance on the system and, by the application ofsuitable data analysis techniques, detailed structural information aboutthe molecular analyte can be obtained.

It is noted that, electrochemical charge transfer rate is a function ofa rate of quantum-mechanical electron transfer from an analyte in anelectrolyte and electrode. In this context, FIG. 4 shows factorsaffecting energy state transition rate in a weakly coupled,non-adiabatic reaction. The transition rate is determined by electronicpolarization difference, electronic coupling energy and energy gap,among other factors. In addition, molecular bonds with vibrations thatare resonant with the energy gap excite electron energy statetransitions that can be measured as leakage current across theelectrochemical interface. Thus, molecular bonds with resonantvibrations can cause leakage currents to flow in addition to backgroundleakage current induced by thermal effects. Thus, the biosensoraccording to various embodiments described herein measures a spectrum ofmolecular vibrational oscillation modes of an analyte at anelectrochemical interface of a liquid electrolyte that is in resonancewith an energy gap limiting the electronic transition potential.

In one sense, this approach has a closer analogy with electromagneticprobes, such as near-infrared (NIR) vibrational spectroscopy, than withconventional electrostatic measurements. However, as molecular structureinformation is transduced directly to an electronic signal beforeacquisition, the biosensor according to the embodiments described hereinis highly scaleable. The direct acquisition of chemistry specificinformation about an analyte in the form of molecular vibrational modesalso eliminates the need for time and labor-intensive combinatorialscreening against bait-molecule probes required by traditionalbioassays.

A quantum information transduction mechanism described herein enableshighly specific interrogation of THz frequency molecular vibrations atexperimentally accessible (˜mV) electronic energies/potentials, byscanning the electronic energy with an applied voltage at a metallicelectrode configure to induce a weak coupling regime. As illustrated inFIG. 5, experimental results suggest sensitivity to a single atom massisotope substitution and sensitivity to structural isomerism. Thissensitivity has not been demonstrated with traditional electronicdetection techniques.

Quantum Tunneling Biosensor Platform

In one embodiment, a thin (about 5-10 mm, for example) dielectric-filmmodified electrode-electrolyte interface is relied upon. The interfacemay be patterned with windows on a silicon die, where theelectrode-electrolyte interface is configured to specifically minimizeelectronic coupling between electrode and electrolyte phases. Leakagecurrent flux is recorded by low or ultra-low noise acquisition circuitryfabricated, for example, by a complementary metal oxide semiconductor(CMOS) process and integrated with the electrode-electrolyte interface.A shielding and interconnect topology is designed to minimize signalcontamination by band-limited noise, electromagnetic interferingsignals, flicker noise, detection artifacts, and any residual de-phasingat the electrode-electrolyte interface. Acquired data may be transmittedfor further filtering, if necessary, as well as for data recording anddisplay.

In other embodiments, the biosensor may further comprise pre-screeningor a means for pre-screening a level of specific biological markersbefore assaying for an analyte of interest. For example, in a biosensortargeting blood toxins, pre-screening for cytokines allows forevaluation of overall health and can indicate presence or absence of abacterial infection.

FIG. 6 shows an exemplary embodiment of a biosensor 60 according to anexample embodiment in which a sensor die with a patternedelectrochemical interface and CMOS interface chip are integrated into alow-cost, disposable, lateral flow-based microfluidic architecture. Inone example use of this embodiment, capillary transport separates serumfrom whole blood and delivers it to the electrode surface. The biosensorplatform comprises elements at the macro-, micro-, and nano-scales,where the microfluidic elements bridge the nano-scale transducer toblood sampling and dispensing at the macro-scale. Since the patternedsensor interface with the integrated electronic is the most expensivecomponent of the platform, the microfluidics is designed such thatfabrication costs are low, power consumption is negligible and themicrofluidic component can be easily disposed of if excessive blockageobstructs the flow path.

In FIG. 6, the biosensor 600 comprises an acquisition zone 602, where asample 604 is dispersed. The sample 604 is then wicked through a modularfluidic system comprising a filtration membrane 606, an immunoseparationmembrane 608, a micro-chromatograph column 610, and an adsorption pad612. With reference to FIG. 7, a disposable module 700 that makes up afluidic system for one embodiment of the biosensor 600 is illustrated. Asample is first wicked through a filtration membrane 606. In thisembodiment, the filtration membrane 606 possesses a graded porestructure capable of separating serum from whole blood. Next the serumpasses through an immunoseparation membrane 608, such as anitrocellulose membrane or other appropriate type of membrane comprisingsurface antibodies specific to high abundance proteins, which remove thehigh-abundance proteins. Finally, the liquid sample moves though themicro-chromatograph column 610, thus fractionating the remainingproteins and results in size separated elutants at the column exit. Themicro-chromatograph column 610 is comprised of a tapered microfluidicchannel containing a photo-polymerized gel.

Although not required for all sample types, a fluidic system module ispreferred when analyzing complex mediums, such as blood, wherecomponents may interfere with the detection of low abundance analytes.Pumping of the sample may be active or passive into the fluidic system.It should be recognized that the filtration media, chosen filtermembranes, other membranes, and characteristics of themicro-chromatography column 610, may be dependent upon factors such assample, abundance of target analyte, and the like.

Referring back to FIG. 6, after the sample 604 passes through themicro-chromatograph column 610, it passes over the adsorption pad 612which provides a thermodynamic gradient for passive capillary actuationof the liquid sample. Liquid fractions eluted from themicro-chromatograph column 610 flow over an active sensor interface 614,which comprises a transducing electrochemical interface integrated withunderlying acquisition electronics, such as CMOS acquisitionelectronics. These components are discussed in further detail withreference to FIGS. 8 and 13 below. A sensor area 616 is patterned aselectrically accessible, thermally insulated, pixilated electrodes forinterrogating the sample 602. It should be appreciated that the sensorpixels 618 can be singular or exist as an array of sensors.

FIG. 8 shows an enlarged view of a sensor package 800 with electrodesensors in arranged in an array. The sensor package 800, in thisembodiment, exists as a layered, heterogeneously integrated sensor-CMOSplatform. In one embodiment, the sensor package 800 comprises a device802, which may be at the exit of the column 610 described with referenceto FIG. 6, to transfer a liquid sample to the electrode sensor array.The device 802 may be embodied as a fluidic chamber and passivation madefrom tubing and epoxy, for example.

The sample then reaches an electrode sensor array 804, which comprises,for example, dielectric thin films deposited by Atomic Layer Depositionand/or Molecular Vapor Deposition (ALD-MVD) on a Si die 805 with anelectrode array and through-silicon vias 811. In one embodiment, theelectrode array 804 is coupled to a CMOS application-specific integratedcircuit (ASIC), which further comprises sensor interface circuitry 806.The sensor interface circuitry 806 is described in further detail withreference to FIG. 13. At the level of sensor interface circuitry 806,the device further comprises electronic pads 808 for connectivity to thebackside of the sensor die 805 and pads 810 for external deviceconnectivity.

Referring back to FIG. 6, the sensor package is mounted on a shieldedprinted circuit board (PCB) for electrical access. Parallel dataacquisition over a large applied bias range is made possible byelectronic-energy-window specific optimization of individual pixels 618,with each pixel nanostructure being optimized for interrogating aspecific electronic energy/bias window. Data acquired as one or moretransition current signals can be transmitted to an external system 650for post-acquisition processing, storage, and display. The biosensorplatform is designed such that sensing and data acquisition modules canbe easily plugged into and out of the fluidic systems, and differentelements, such as elements 602, 606, 608, 610, 612, and 614 of thefluidic system exist as modules, so that the platform can be dissembled,interchanged, and disposed of if necessary.

In one embodiment, resolved spectral information, once acquired, is thencorrelated with vibrational energy data to identify specific molecularspecies associated with a macro-molecule analyte. This may beaccomplished by employing an information-driven strategy for targeted,non-redundant analysis of bio-analyte(s) in an electrolyte solution.Signatures of information-rich subsets of a bio-analyte, such ascysteine-containing peptides, phosphorylated peptides, or glycosylatedpeptides, can be tracked in the resolved spectrum of the bio-analyte.These subsets can serve as molecular markers for identifying andquantifying the presence of the molecular species of interest. Areference database containing these molecular markers can be constructedfor each target analyte. In other words, each analyte can produce itsown signature spectrum of these information rich subsets. By comparingthe resolved spectrum from the sample to the reference database, it ispossible to identify the target analyte.

FIG. 9 shows steps in generating a reference database. At referencenumeral 902, a purified recombinant form of the target molecule (orbiological surrogate, in the case of neurotoxins) is systematicallydigested by enzymes, such as trypsin and chymotrypsin, for example, togenerate peptide fragments. This is followed by a three-dimensionalseparation technique at reference numeral 904, such as 2-D polyacrylamide gel electrophoresis (2-D PAGE) in tandem with highperformance liquid chromatography (HPLC), preferably reverse phase-HPLC.Fractions are then collected, purified, and re-extracted in a suitablebuffer and analyzed using embodiments of the biosensor or quantumtunneling electronic biosensor described herein at reference numeral906. The data, after background subtraction, is analyzed forcharacteristic spectra of moieties specific to the peptide fragment inthe aliquot being tested at reference numeral 908. Stable isotopes ofreference peptides, for example, may be prepared at reference numeral910. In some embodiments, the same fractions as well as theisotope-labeled reference peptides prepared at reference numeral 910 mayalso be examined in parallel by traditional HPLC-MS-MS techniques atreference numeral 912.

Nanoscale Design of the Electrochemical Interface

For some embodiments, electronic de-coupling is enhanced by configuringthe electrochemical interface to minimize overlap between a metalelectrode 1001 and electrolyte phase electronic energy surfaces whilesimultaneously scaling down the transducing surface area. FIG. 10 showstunneling barriers ϕ₁ and ϕ₂ at metal-dielectric 1002 anddielectric-electrolyte 1004 interfaces. Coupling between initial andfinal electronic energy states in the weakly coupled, non-adiabatictransition is modulated by the electrostatic tunneling barrier ϕ₂located at the dielectric-electrolyte interface 1004 as well as by aneffective barrier limiting charge injection at the electrode-dielectricinterface 1002. Nanoscale engineering of barrier heights at theelectrode-dielectric thin-film interface 1002 as well as at thedielectric thin-film-electrolyte interface 1004 is utilized to minimizeenvironmental de-phasing due to electronic coupling.

The desired minimization may be achieved by increasing the electrontunneling barrier ϕ₂ at the dielectric-electrolyte interface. In oneembodiment, the electron tunneling barrier ϕ₂ is increased by increasingthe electrolyte pH. Other exemplary embodiments achieve an increasedtunneling barrier by increasing electrolyte anion electronegativity,increasing dielectric monolayer functional group electronegative, and/orincreasing dielectric monolayer thickness, for example.

The desired minimization may also be attained by increasing the limitingbarrier ϕ₂ at the dielectric-electrolyte interface 1004. In oneembodiment this is achieved by coating the dielectric monolayer 1003with an organic coating, such as short chain silanes, with differentelectronegative, electrolyte-facing functional groups, such as —OH, —OR,—COOH, —SH, —SR, —COR, —NO₂, —Br, and the like. These aforementionedcoatings are suitable for forming with molecular-vapor-deposition at thedielectric surface 1003.

The metal electrode-dielectric barrier 1002, unlike thedielectic-electrolyte barrier 1004, is a function of the metalwork-function, dielectric band gap, and nature of molecular orbitaldistortion induced by a bond between the metal 1001 and dielectricmaterials 1003. A reduction in the coupling between the electrode 1001and the electrolyte 1005 may also be achieved by altering the tunnelingbarrier ϕ₁ located at the dielectric-electrode interface. For example,increasing the tunneling barrier ϕ₁ at the dielectric-electrode surfacereduces coupling between electrode and electrolyte phases, thus leadingto increased resolution of vibrational frequency information in themeasured current. For some embodiments, the mechanism of tunneling basedcharge injection in the dielectric 1003 would be electron tunneling. Inother words, the dielectric 1003 would be comprised of aninorganic-oxide. For these embodiments, metals such as Pt, Ir, Se, or Auare preferred. For other embodiments, hole-tunneling is the mechanism ofcharge injection in the dielectric 1003. One having an ordinary skillwould appreciate that the dielectric film 1003 in these embodimentswould be comprised of an organic alkane. For these embodiments, metalslike Ta, Ti, Zr, or Hf are preferred. The final choice of metal isdependent on many factors including the mechanical, diffusional, andelectrochemical stability of the electrode, ease of deposition,electrical resistivity, and ability to seed a dielectric layer, forexample.

In various embodiments, the dielectric film 1003 that spatiallyseparates the electrode 1001 and electrolyte 1005 layers is comprised ofa high-k, nanolaminate. The high-k material may be Ta₂O₂, ZrO₂, TiO₂, orother suitable material. It is noted that large dielectric constants forthe insulating film facilitate greater charge accumulation at thedielectric-electrolyte interface, thus effectively increasing thetunneling barrier and reducing coupling. However, a larger dielectricconstant is typically associated with a small band-gap and,consequently, higher non-tunneling leakage current. Thus, the high-kmaterial may be intercalated between alternating layers oflower-dielectric constant oxides, such as c or other suitable compoundor compounds. In this context, FIG. 11 shows sequential layering of high1102 and low 1104 k-dielectric materials for a high-k dielectricinsulator. The final choice of materials used to form a given dielectricnanolaminate is determined by insulator properties like breakdownresistance, electrochemical and mechanical stability, and chemicalinertness to aqueous electrolytes in the presence of an applied bias,for example.

Alternatively to mechanical alterations, a reduction in energy statecoupling may be achieved by applying a directional magnetic field to theelectrochemical interface. In this case, a reduction in coupling mayoccur by interaction of the magnetic field with a magnetic dipole momentof a transitioning electron, similar to the interaction effect betweenthe applied electric field and charge of the transitioning electron. Invarious embodiments, a nanolaminate for a dieletric film that reliesupon magnetic tunneling uses differentially oriented film magneticmoments to further restrict electronic transition. Preferably, amagnetic tunneling nanolaminate can comprise dieletric-based thin-filmarchitectures with room temperature ferromagnetic properties that allowfor the generation of local, inhomogeneous magnetic fields that caninteract with the magnetic dipole of the transitioning electron to“gate” the quantum-mechanical transition. In this context, FIG. 12 showsa magnetic tunneling film architecture that uses differentially orientedfilm magnetic moments to further restrict electronic transition. Amulti-stage-gate-like design of the dielectric nanolaminate may enablefurther minimization of the de-phasing resulting from electroniccoupling between the initial and final energy states. In one embodiment,the dielectric film comprises low-k (e.g., HfO₂) 1204/high-k 1202dielectric substacks interspersed with thin films of a non-magneticdielectric insulator 1206 with an intermediate value of dielectricpermittivity, such as Al₂O₃. Aluminum in the Al₂O₃ thin-film lack thed-shell orbitals necessary for displaying magnetic susceptibility andthus the alumina thin-films are believed to be non-magnetic, making themsuited for this application. In other words, the functionalferromagnetic elements of the dielectric thin film is made up of low-k(e.g., HfO₂) 1204/high-k 1202 dielectric substacks and every twosubs-stacks are separated by Al₂O₃ thin film which functions as aninsulating barrier that minimizes conservative and dissipative magneticcoupling between adjoining ferromagnetic sub-stacks. The total number ofsubstacks and Al₂O₃ thin films and, thus, the extent offerromagnetic-induced decoupling, is limited by the overall thickness ofthe dielectric film, which is less than approximately 10 mm in variousembodiments.

Analog Front-End Instrumentation

A key component to the biosensor according to embodiments describedherein is analog front-end instrumentation, which is relied upon toachieve minimum noise in an acquired signal. Physical and electronicsources of noise in a measured transition rate limit the extent ofdecoupling achievable and can become measurement-limiting whentransition rates in the weakly-coupled regime are on the order of 10 fA.As stated earlier, scaling down the active electrode area supplementselectronic decoupling achieved by engineering the nano-scale,information transducing interface, and simultaneously miniaturizes thesensing platform. Similarly, noise in the system also limits the extentto which the electrode area can be minimized. In sum, to achieveeffective device minimization, rapid analyte identification, and batterylife preservation, non-electron transfer leakage currents and noise inthe acquisition electronics should preferably be minimized.

A measured non-adiabatic current is a function of two non-interactingfrequency domains: (a) a “macro-frequency” (approximately 1 Hz) thatdetermines the rate-limiting step in the macroscopic electrochemicalsystem and (b) a “micro frequency” (>10¹² Hz) that measures the dynamicsof molecular vibrations, where the dynamics are manifest in theelectronic energy, or applied bias, space. A low-frequency AlternatingCurrent (AC) excitation is applied to the electrochemical interface, andthe current is recorded as a function of Direct Current (DC) bias at theelectrode. The non-adiabatic transition rate is contaminated, at leastin part, by band-limited white noise and electromagnetic interference.Effective signal extraction requires suppression of both extrinsic andintrinsic noise contributions. In short, because of the low frequencytransition signal strength, a high signal-to-noise ratio is preferable.

FIG. 13 shows an example of three-electrode feedback suppression ofthermal noise for electronic transition measurements. In one embodiment,a three-electrode analog measurement topology circuit 1300 is used forhigh gain feedback suppression of voltage-noise in the macro-frequencydomain. The circuit 1300 includes a first gain amplifier 1302, a secondgain amplifier 1304 coupled to an output of the first gain amplifier1302, and a counter electrode 1306 coupled to an output of the secondgain amplifier 1304. The circuit 1300 further includes a referenceelectrode 1308 and a working electrode 1310, each coupled to a biasamplifier 1312. An output of the bias amplifier 1312 is coupled to a DCgenerator 1314. Outputs of the DC generator 1314 and an AC generator1316 are combined at a node 1317 in the circuit 1300 and provided as aninput to the second gain amplifier 1304. As illustrated in FIG. 13, thecombined outputs of the DC and AC generators 1314 and 1316 are furthercombined with the output of the first gain amplifier 1302 at the node1317 before being provided as input to the second gain amplifier 1304.

The circuit 1300 further includes a current measurement circuit 1318coupled to the working electrode 1310, and a lock-in detection circuit1320 coupled to an output of the current measurement circuit 1318. Asillustrated in FIG. 13, a reference signal output of the lock-indetection circuit 1320 is provided as an input to the AC generator 1316,an out-of-phase signal output of the lock-in detection circuit 1320 isprovided as an input to the first gain amplifier 1302, and an in-phasesignal output of the lock-in detection circuit 1320 is provided as aninput to a data acquisition element 1322 for processing.

According to one embodiment, a feedback loop of the circuit 1300 acts tominimize thermal noise on the order of 40 pA/Hz¹¹², at room temperature.In embodiments that have tunneling transition currents on the order of10 fA, noise on the order of 2 fA rms/Hz¹¹² would be sufficient for theresolution of spectral features in measured data. In this context, forexample, FIG. 14 shows a table containing program metrics.

Sensing of the low frequency modulated energy state transition signalmay be performed with low noise current amplification and readoutcircuitry. The signal acquisition path employs chopper modulation tomitigate flicker noise from CMOS devices, which may be expected to besignificant around the “macro” frequency modulation. In one embodiment,a metal-oxide-semiconductor-field-effect transistor (MOSFET) thatcomprises defect free or substantially defect free oxide interfacesalong a channel region is relied upon. This eliminates the trapping anddetrapping of carriers that create 1/f noise characteristics.

Although exemplary embodiments have been shown and described, it will beclear to those of ordinary skill in the art that a number of changes,modifications, or alterations to the disclosure as described may bemade.

At least the following is claimed:
 1. A tunneling biosensor interfacesystem for sensing chemical information, the system comprising: afluidic system comprising: a sample acquisition zone configured toreceive a sample comprising a redox specie and an analyte specie; afiltration module coupled to the sample acquisition zone and configuredto wick the sample from the sample acquisition zone and separate theanalyte specie from the sample; and a quantum tunneling biosensorinterface coupled to the fluidic system and configured to detect theanalyte specie, the quantum tunneling biosensor interface comprising: atransducing electrode array comprising dielectric thin films depositedon an electrode array wherein the dielectric thin films comprise ananolaminate to apply a directional magnetic field across thetransducing electrode array; and sensor interface circuitry coupled tothe transducing electrode array and configured to detect a tunnelingcurrent configured to flow between the redox specie and the transducingelectrode array via at least one dielectric thin film deposited on anelectrode array, wherein the tunneling current is indicative of theanalyte specie.
 2. The system of claim 1, further comprising a voltagesource to apply a voltage bias across the transducing electrode array toproduce the tunneling current.
 3. The system of claim 1, wherein thequantum tunneling biosensor interface is mounted on a shielded printedcircuit board.
 4. The system of claim 1, wherein the sensor interfacecircuitry is coupled to the transducing electrode array bythrough-silicon vias in a silicon die.
 5. The system of claim 1, whereinthe filtration module includes a graded pore structure configured toseparate the analyte specie from the sample.
 6. The system of claim 1,wherein the dielectric thin films are configured to spatially separatethe sample and the sensor interface circuitry.
 7. The system of claim 1,wherein the fluidic system further comprises: an immunoseparation modulecoupled to the filtration module; a tapered micro-chromatograph coupledto the immunoseparation module; and an adsorption pad coupled to thequantum tunneling biosensor interface.
 8. The system of claim 7, whereinthe immunoseparation module includes a nitrocellulose membrane includingsurface antibodies configured to remove high-abundance proteins from thesample.
 9. The system of claim 1, wherein the dielectric thin filmscomprise at least one nanolaminate having a high dielectric constant.10. The system of claim 9, wherein the at least one nanolaminatecomprises high dielectric constant layers and low dielectric constantlayers, the high dielectric constant layers being intercalated betweenthe low dielectric constant layers.
 11. The system of claim 10, whereinthe high dielectric constant layers comprise at least one materialselected from the group consisting of: HfO₂; Ta₂O₂; ZrO₂; and TiO₂. 12.The system of claim 10, wherein the low dielectric constant layerscomprise at least one organic alkane layer.
 13. The system of claim 1,wherein the dielectric thin films comprising layers of a non-magneticdielectric insulator intercalated between substacks, the substackscomprising alternating layers of a first ferromagnetic material with ahigh dielectric constant and a second ferromagnetic material with a lowdielectric constant.
 14. The system of claim 13, the non-magneticdielectric insulator comprising Al₂O₃.
 15. A tunneling interface,comprising: a fluidic system comprising: a sample acquisition zoneconfigured to receive a sample comprising a redox specie and an analytespecie; and a filtration module coupled to the sample acquisition zoneand configured to wick the sample from the sample acquisition zone andseparate the analyte specie from the sample; a quantum tunnelingbiosensor interface coupled to the fluidic system and configured todetect the analyte specie, the quantum tunneling biosensor interfacecomprising: a transducing electrode sensor array comprising at least onedielectric thin film layered on an electrode array wherein the at leastone dielectric thin film comprises a nanolaminate to apply a directionalmagnetic field across the transducing electrode array; a voltage sourceto apply a voltage bias across the transducing electrode sensor array toproduce a tunneling current between the redox specie and the transducingelectrode array via the at least one dielectric thin film deposited onan electrode array; and sensor interface circuitry coupled to thetransducing electrode sensor array, and configured to detect thetunneling current, wherein the tunneling current is indicative of theanalyte specie.
 16. The tunneling interface of claim 15, furthercomprising: an immunoseparation module in fluid communication with thesample acquisition zone; and a tapered micro-chromatograph in fluidcommunication with the immunoseparation module and the transducingelectrode sensor array.
 17. The tunneling interface of claim 15, whereinthe transducing electrode sensor array comprises the at least onedielectric thin film layered on a metal electrode array.
 18. Thetunneling interface of claim 15, wherein the at least one dielectricthin film comprises a high dielectric constant layer comprising thenanolaminate.
 19. The tunneling interface of claim 15, wherein the atleast one dielectric thin film comprises at least one high dielectricconstant layer and at least one low dielectric constant layer.
 20. Asystem for sensing chemical information, the system comprising: afluidic system, comprising: a sample acquisition zone; a taperedmicro-chromatogram operatively coupled to the sample acquisition zone;and an adsorption pad operatively coupled to the taperedmicro-chromatogram; and a quantum tunneling biosensor interfaceoperatively coupled to the adsorption pad, the quantum tunnelingbiosensor interface comprising: a transducing electrode array comprisingdielectric thin films deposited on an electrode array, the dielectricthin films being configured to produce a weakly-coupled non-adiabaticelectron flux in response to a voltage bias being applied to thetransducing electrode array; and sensor interface circuitry processinglogic operatively coupled to the transducing electrode array.
 21. Thesystem of claim 20, further comprising a voltage source to apply thevoltage bias across the transducing electrode array to produce theweakly-coupled non-adiabatic electron flux.
 22. The system of claim 20,wherein the dielectric thin films comprise at least one high dielectricconstant layer and at least one low dielectric constant layer.